U.S. patent application number 11/521960 was filed with the patent office on 2007-05-10 for pulse oximetry signal correction using near infrared absorption by water.
Invention is credited to Clark R. JR. Baker, Carine Hoarau, Edward Karst.
Application Number | 20070106137 11/521960 |
Document ID | / |
Family ID | 34920066 |
Filed Date | 2007-05-10 |
United States Patent
Application |
20070106137 |
Kind Code |
A1 |
Baker; Clark R. JR. ; et
al. |
May 10, 2007 |
Pulse oximetry signal correction using near infrared absorption by
water
Abstract
A method and system for measuring a physiological parameter,
comprising collecting a first absorbance at a first wavelength,
chosen to be primarily absorbed by water; collecting a second
absorbance at a second wavelength, chosen to be primarily absorbed
by hemoglobin; and combining the first signal and the second signal
to generate a combined plethysmograph signal which is proportionate
lower in noise caused by motion-related interference.
Inventors: |
Baker; Clark R. JR.;
(Newman, CA) ; Karst; Edward; (South Pasadena,
CA) ; Hoarau; Carine; (Lafayette, CA) |
Correspondence
Address: |
FLETCHER YODER (TYCO INTERNATIONAL, LTD.)
P.O. BOX 692289
HOUSTON
TX
77269-2289
US
|
Family ID: |
34920066 |
Appl. No.: |
11/521960 |
Filed: |
September 15, 2006 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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10797475 |
Mar 9, 2004 |
|
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11521960 |
Sep 15, 2006 |
|
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Current U.S.
Class: |
600/336 |
Current CPC
Class: |
A61B 5/14551 20130101;
A61B 5/7207 20130101; A61B 5/02416 20130101 |
Class at
Publication: |
600/336 |
International
Class: |
A61B 5/00 20060101
A61B005/00 |
Claims
1. A method of determining a physiological parameter, comprising:
obtaining a first absorbance at a first wavelength, wherein the
first wavelength is chosen to be primarily absorbed by water;
obtaining a second absorbance at a second wavelength, wherein the
second wavelength is chosen to be primarily absorbed by hemoglobin;
estimating a first ratio-of-ratios between the first absorbance and
the second absorbance; and using the first ratio-of-ratios to
identify motion noise.
2. The method of claim 1, comprising obtaining a third absorbance
at a third wavelength, wherein the third wavelength is chosen to be
primarily absorbed by hemoglobin; calculating a corrected second
absorbance by subtracting a first fraction of the first absorbance
from the second absorbance; and calculating a corrected third
absorbance by subtracting a second fraction of the first absorbance
from the third absorbance.
3. The method of claim 2, comprising: resealing the corrected
second absorbance to obtain a rescaled second absorbance; and
rescaling the corrected third absorbance to obtain a rescaled third
absorbance, wherein the rescaling of the corrected second
absorbance and the corrected third absorbance maintains a
ratio-of-ratios between the rescaled second absorbance and the
rescaled third absorbance that is independent of changes in the
subtracted first and section fractions of the first absorbance.
4. The method of claim 3, further comprising: calculating a first
oxygen saturation value using a two wavelength algorithm based on
the rescaled second absorbance and the rescaled third absorbance;
and calculating a second oxygen saturation value using a two
wavelength algorithm based on the second absorbance and the third
absorbance.
5. The method of claim 3, further comprising: calculating an oxygen
saturation value using a two wavelength algorithm based on the
rescaled second absorbance and the rescaled third absorbance.
6. The method of claim 4, further comprising: displaying the first
oxygen saturation value during periods of high motion noise;
displaying the second oxygen saturation value during periods of low
motion noise; calculating an intermediate oxygen saturation value
using the first and second oxygen saturation values; and displaying
the intermediate oxygen saturation value during periods of
intermediate motion noise.
7. The method of claim 3, further comprising: calculating a first
pulse rate value using the rescaled second absorbance; and
calculating a second pulse rate value using the second
absorbance.
8. The method of claim 3, further comprising: calculating a pulse
value using the corrected second absorbance or the rescaled second
absorbance.
9. The method of claim 7, further comprising: displaying the first
pulse rate value during periods of high motion noise; displaying
the second pulse rate value during periods of low motion noise;
calculating an intermediate pulse rate value using the first and
second pulse rate values; and displaying the intermediate pulse
rate value during periods of intermediate motion noise.
10. The method of claim 2, further comprising using an algorithm to
calculate the value of the first fraction and the value of the
second fraction.
11. The method of claim 10, wherein the algorithm minimizes the
standard deviation of a resulting waveform.
12. The method of claim 10, wherein the algorithm minimizes the
power of a resulting waveform.
13. The method of claim 10, wherein the algorithm minimizes the
skewness of a resulting waveform.
14. The method of claim 10, wherein the algorithm is chosen to
reduce errors in the calculation of oxygen saturation and pulse
rate.
15. The method of claim 10, wherein the algorithm comprises an
infinite impulse response (IIR) filter.
16. A system for signal correction in pulse oximetry, the system
comprising: a pulse oximeter monitor, wherein the pulse oximeter
monitor is configured to analyze a first absorbance signal
substantially corresponding to motion noise, to analyze a second
absorbance signal chosen to be primarily absorbed by hemoglobin, to
analyze a third absorbance signal chosen to be primarily absorbed
by hemoglobin, to adjust the second absorbance signal to compensate
for noise by subtracting a first fraction of the first absorbance
signal from the second absorbance signal to obtain a corrected
second absorbance signal, and to adjust the third absorbance signal
to compensate for noise by subtracting a second fraction of the
first absorbance signal from the third absorbance signal to obtain
a corrected third absorbance signal.
17. The system of claim 16, further comprising a pulse oximetry
sensor, wherein the sensor comprises optical emitters and detectors
configured to emit and detect light at a first wavelength chosen to
be primarily absorbed by water, to emit and detect light at a
second wavelength chosen to be primarily absorbed by hemoglobin,
and to emit and detect light at a third wavelength chosen to be
primarily absorbed by hemoglobin.
18. The system of claim 16, wherein the pulse oximeter monitor is
configured to combine the first absorbance signal and the second
absorbance signal to obtain a metric that identifies the presence
of motion noise.
19. The system of claim 16, wherein the pulse oximeter monitor is
configured to rescale the corrected second absorbance signal to
obtain a rescaled second absorbance signal, and to rescale the
corrected third absorbance signal to obtain a rescaled third
absorbance signal, wherein the resealing maintains a
ratio-of-ratios between the rescaled second absorbance signal and
the rescaled third absorbance signal that is independent of changes
in the subtracted first and section fractions of the first
absorbance.
20. The system of claim 19, wherein the pulse oximeter monitor is
configured to calculate a first oxygen saturation value using the
rescaled second absorbance signal and the rescaled third absorbance
signal, and to calculate a second oxygen saturation value using the
first absorbance signal and the second absorbance signal.
21. The system of claim 19, wherein the pulse oximeter monitor is
configured to calculate an oxygen saturation value using the
rescaled second absorbance signal and the rescaled third absorbance
signal.
22. The system of claim 20, wherein the pulse oximeter monitor is
configured to display the first oxygen saturation value during
periods of high motion noise, to display the second oxygen
saturation value during periods of low motion noise, to calculate
an intermediate oxygen saturation value using the first oxygen
saturation value and the second oxygen saturation value, and to
display the intermediate oxygen saturation value during periods of
intermediate motion noise.
23. The system of claim 19, wherein the pulse oximeter monitor is
configured to calculate a first pulse rate value from the rescaled
second absorbance signal and the rescaled third absorbance signal,
and to calculate a second pulse rate value from the second
absorbance signal and the third absorbance signal.
24. The system of claim 19, wherein the pulse oximeter monitor is
configured to calculate a pulse rate value from the rescaled or
corrected second and third absorbance signals.
25. The system of claim 23, wherein the pulse oximeter monitor is
configured to display the first pulse rate value during periods of
high motion noise, to display the second pulse rate value during
periods of low motion noise, to calculate an intermediate pulse
rate value using the first pulse rate value and the second pulse
rate value, and to display the intermediate pulse rate value during
periods of intermediate motion noise.
26. One or more tangible, machine readable media, comprising code
executable to perform the acts of: obtaining a first absorbance at
a first wavelength, wherein the first wavelength is chosen to be
primarily absorbed by water; obtaining a second absorbance at a
second wavelength, wherein the second wavelength is chosen to be
primarily absorbed by hemoglobin; estimating a first
ratio-of-ratios between the first absorbance and the second
absorbance; and using the first ratio-of-ratios to identify motion
noise.
27. The one or more tangible, machine readable media of claim 26,
further comprising code executable to perform the acts of:
obtaining a third absorbance at a third wavelength, wherein the
third wavelength is chosen to be primarily absorbed by hemoglobin;
calculating a corrected second absorbance by subtracting a first
fraction of the first absorbance from the second absorbance; and
calculating a corrected third absorbance by subtracting a second
fraction of the first absorbance from the third absorbance.
28. The one or more tangible, machine readable media of claim 27,
further comprising code executable to perform the acts of:
rescaling the corrected second absorbance to obtain a rescaled
second absorbance; and rescaling the corrected third absorbance to
obtain a rescaled third absorbance, wherein the rescaling of the
corrected second absorbance and the corrected third absorbance
maintains a ratio-of-ratios between the rescaled second absorbance
and the rescaled third absorbance that is independent of changes in
the subtracted first and section fractions of the first
absorbance.
29. The one or more tangible, machine readable media of claim 28,
further comprising code executable to perform the acts of:
calculating a first oxygen saturation value using a two wavelength
algorithm based on the rescaled second absorbance and the rescaled
third absorbance; and calculating a second oxygen saturation value
using a two wavelength algorithm based on the second absorbance and
the third absorbance.
30. The one or more tangible, machine readable media of claim 27,
further comprising code executable to perform the acts of:
Calculating an oxygen saturation value using a two wavelength
algorithm based on the rescaled second absorbance and the rescaled
third absorbance.
31. The one or more tangible, machine readable media of claim 29,
further comprising code executable to perform the acts of:
displaying the first oxygen saturation value during periods of high
motion noise; displaying the second oxygen saturation value during
periods of low motion noise; calculating an intermediate oxygen
saturation value using the first oxygen saturation value and the
second oxygen saturation value; and displaying the intermediate
oxygen saturation value during periods of intermediate motion
noise.
32. The one or more tangible, machine readable media of claim 28,
further comprising code executable to perform the acts of:
calculating a first pulse rate value using the rescaled second
absorbance; and calculating a second pulse rate value using the
second absorbance.
33. The one or more tangible, machine readable media of claim 28,
further comprising code executable to perform the acts of:
calculating a pulse rate value using the rescaled second absorbance
or the corrected second absorbance.
34. The one or more tangible, machine readable media of claim 32,
further comprising code executable to perform the acts of:
displaying the first pulse rate value during periods of high motion
noise; displaying the second pulse rate value during periods of low
motion noise; calculating an intermediate pulse rate value using
the first pulse rate value and the second pulse rate value; and
displaying the intermediate pulse rate value during periods of
intermediate motion noise.
35. The one or more tangible, machine readable media of claim 27,
further comprising code executable to perform the act of using an
algorithm to calculate the value of the first fraction and the
value of the second fraction.
36. The one or more tangible, machine readable media of claim 35,
wherein the algorithm minimizes the standard deviation of a
resulting waveform.
37. The one or more tangible, machine readable media of claim 35,
wherein the algorithm minimizes the power of a resulting
waveform.
38. The one or more tangible, machine readable media of claim 35,
wherein the algorithm minimizes the skewness of a resulting
waveform.
39. The one or more tangible, machine readable media of claim 35,
wherein the algorithm is chosen to reduce errors in the calculation
of oxygen saturation and pulse rate.
40. The one or more tangible, machine readable media of claim 35,
wherein the algorithm is an infinite impulse response (IIR) filter.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation-in-part of the U.S.
patent Ser. No. 10/797,475, entitled "PULSE OXIMETRY MOTION
ARTIFACT REJECTION USING NEAR INFRARED ABSORPTION BY WATER", filed
Mar. 9, 2004, which is herein incorporated by reference in its
entirety.
BACKGROUND OF THE INVENTION
[0002] 1. Field of the Invention
[0003] The present invention relates generally to the processing of
signals obtained from a medical diagnostic apparatus, such a pulse
oximeter, using near infrared spectroscopy, to remove artifact or
noise effects from the signal representative of a physiological
parameter of interest.
[0004] 2. Description of the Related Art
[0005] This section is intended to introduce the reader to various
aspects of art that may be related to various aspects of the
present invention, which are described and/or claimed below. This
discussion is believed to be helpful in providing the reader with
background information to facilitate a better understanding of the
various aspects of the present invention. Accordingly, it should be
understood that these statements are to be read in this light, and
not as admissions of prior art.
[0006] A typical pulse oximeter measures two physiological
parameters, percent oxygen saturation of arterial blood hemoglobin
(SpO.sub.2 or sat) and pulse rate. Oxygen saturation can be
estimated using various techniques. In one common technique, the
photocurrent generated by the photo-detector is conditioned and
processed to determine the ratio of modulation ratios (ratio of
ratios) of the red to infrared signals. This modulation ratio has
been observed to correlate well to arterial oxygen saturation. The
pulse oximeters and sensors are empirically calibrated by measuring
the modulation ratio over a range of in vivo measured arterial
oxygen saturations (SaO.sub.2) on a set of patients, healthy
volunteers, or animals. The observed correlation is used in an
inverse manner to estimate blood oxygen saturation (SpO.sub.2)
based on the measured value of modulation ratios of a patient. Most
pulse oximeters extract the plethysmographic signal having first
determined saturation or pulse rate, both of which are susceptible
to interference.
[0007] In general, pulse oximetry takes advantage of the fact that
in live human tissue, hemoglobin is a strong absorber of light
between the wavelengths of 500 and 1100 nm. The pulsation of
arterial blood through tissue is readily measurable, using light
absorption by hemoglobin in this wavelength range. A graph of the
arterial pulsation waveform as a function of time is referred to as
an optical plethysmograph. The amplitude of the plethysmographic
waveform varies as a function of the wavelength of the light used
to measure it, as determined by the absorption properties of the
blood pulsing through the arteries. By combining plethysmographic
measurements at two different wavelength regions, where oxy- and
deoxy-hemoglobin have different absorption coefficients, the oxygen
saturation of arterial blood can be estimated. Typical wavelengths
employed in commercial pulse oximeters are 660 and 890 nm.
[0008] It is known that rapid motion or application of pressure to
a tissue site can have the effect of changing the optical
properties being measured at or near that site. The amplitude of
the optical signal changes associated with such events, known as
motion artifacts, can easily be larger than that due to the
arterial pulse. In practice, this can lead to inaccurate estimation
of the percent oxygen saturation by pulse oximetry. Various
techniques for addressing and removing undesired signal effects,
including motion artifacts are known. As used herein, noise refers
to signal portions that are undesired or are not directly related
to changes in optical properties that are related to the arterial
pulse, and which may include motion artifact. The optical signal
through the tissue can be degraded by both noise and motion
artifact. One source of noise is ambient light which reaches the
light detector. Another source of noise is electromagnetic coupling
from other electronic instruments. Motion of the patient also
introduces noise and affects the signal. For example, the contact
between the detector and the skin, or the emitter and the skin, can
be temporarily disrupted when motion causes either to move away
from the skin. In addition, since blood is a fluid, it responds
differently than the surrounding tissue to inertial effects, thus
resulting in momentary changes in volume at the point near which
the oximeter probe is attached.
[0009] Motion artifact can degrade a pulse oximetry signal relied
upon by a health care provider, without the provider's awareness.
This is especially true if the monitoring of the patient is remote,
the motion is too small to be observed, or the health care provider
is watching the instrument or other parts of the patient, and not
the sensor site. There are various known techniques for addressing
the effects of noise and/or motion artifacts.
[0010] For example, U.S. Pat. No. 4,714,341 discloses a method for
combining three wavelengths to detect the presence of motion. The
wavelengths are used two at a time to separately compute the oxygen
saturation percentage. When the oxygen saturation values computed
using different wavelength combinations are in poor agreement, this
is assumed to be caused by motion artifact, and the value is
discarded. A disadvantage of this approach is that the agreement or
lack thereof between the saturation values may or may not be due to
motion artifact. In addition, this approach does not identify or
remove the effects of motion artifact, but instead discards values
that appear suspect
[0011] Another approach involves the filtering of pulse oximetry
signals. However, filtering methods require assumptions about the
properties of the artifact that do not always hold in practice. In
addition, this approach does not measure the motion-induced
signal.
[0012] U.S. Pat. No. 5,482,036 provides another approach, and
describes a signal processing method for artifact reduction that
functions when the artifact-related signal is associated with blood
that is at a lower oxygen saturation than the arterial blood. Such
a method relies on the generation of an artificial noise signal,
which is combined with the physiological parameter to reduce the
effect of the unknown noise signal. This approach for reducing the
effects of artifact, without separately measuring the motion
signal, is based on assumptions about the effect of motion on the
plethysmographic signal. Assumptions may or may not be true, and
many assumptions are invalid
[0013] Each of the known techniques for compensating for motion
artifact has its own limitations and drawbacks. It is therefore
desirable that a pulse oximetry system be designed which more
effectively and accurately reports blood-oxygen levels during
periods of motion. While many have attempted to isolate the effects
of undesired signal portions, such as motion-induced artifacts, by
making potentially invalid assumptions or by rejecting suspect
estimates of desired signal values, there still remains a need for
a deterministic identification, determination and measurement of
artifact signals, to enable an accurate measurement of the desired
signal values in the presence of undesired signal portions.
SUMMARY
[0014] Certain aspects commensurate in scope with the originally
claimed invention are set forth below. It should be understood that
these aspects are presented merely to provide the reader with a
brief summary of certain forms of the invention might take and that
these aspects are not intended to limit the scope of the invention.
Indeed, the invention may encompass a variety of aspects that may
not be set forth below.
[0015] By measuring the artifact signal, the present technique
allows motion artifacts to be separated from the plethysmographic
signal without the limiting assumptions of prior known techniques.
The present technique provides methods for measuring the motion
signal associated with changes in tissue optical properties and
using the measurement to compensate plethysmographic measurements
made at other wavelengths.
[0016] In one embodiment, the present technique provides a method
of determining a physiological parameter, including measuring an
absorbance at a wavelength chosen to be primarily absorbed by
water, and measuring an absorbance at a wavelength chosen to be
primarily absorbed by hemoglobin. A ratio-of-ratios is calculated
between these absorbances, and the ratio-of-ratios is used to
identify motion noise.
[0017] In another embodiment, there is provided a system for the
minimization of motion noise artifacts in pulse oximetry. This
system uses a pulse oximeter monitor configured to analyze an
absorbance signal that is primarily reflective of motion noise, an
absorbance signal chosen to be primarily absorbed by hemoglobin,
and another absorbance signal chosen to be primarily absorbed by
hemoglobin. In another aspect, the monitor may be configured to use
the first and second absorbances to obtain a metric that identifies
the presence of motion noise.
[0018] In another embodiment, there is provided one or more
tangible, machine readable media, containing code which controls
the measurement of an absorbance at a wavelength chosen to be
primarily absorbed by water and the measurement of an absorbance at
a wavelength chosen to be primarily absorbed by hemoglobin. Under
the control of this code, a ratio-of-ratios is calculated between
these absorbances, and the ratio-of-ratios is used to identify
motion noise.
BRIEF DESCRIPTION OF THE DRAWINGS
[0019] Advantages of the invention may become apparent upon reading
the following detailed description and upon reference to the
drawings in which:
[0020] FIG. 1 is a block diagram of an exemplary oximeter, in
accordance with aspects of the present technique.
[0021] FIG. 2 is a graph of the plethysmographic amplitude measured
on the human ear as a function of wavelength.
[0022] FIG. 3 is a graph of absorption spectra of the principal
components in human blood, scaled to typical physiological
concentration.
[0023] FIG. 4 is a graph of absorption spectra of the principal
components in human skin, scaled to typical physiological
concentration.
[0024] FIG. 5 is a graph of absorption spectra of the principal
components in human skin, scaled to equal volume-fraction
concentration.
[0025] FIG. 6 is a graph of plethysmographs measured on a human ear
at 4 different wavelengths of approximately 920, 1050, 1180 and
1300 nm respectively.
[0026] FIG. 7 is a graph of an exemplary plethysmographic artifact
reduction by combining measurements at 2 near infrared
wavelengths.
[0027] FIG. 8 is a flowchart of one approach to using a third
wavelength to compensate for motion artifacts, in accordance with
aspects of the present technique.
[0028] FIG. 9 is a graph of the oximetry results showing the error
compensation using the recombination technique.
DETAILED DESCRIPTION OF SPECIFIC EMBODIMENTS
[0029] One or more specific embodiments of the present invention
will be described below. In an effort to provide a concise
description of these embodiments, not all features of an actual
implementation are described in the specification. It should be
appreciated that in the development of any such actual
implementation, as in any engineering or design project, numerous
implementation-specific decisions must be made to achieve the
developers' specific goals, such as compliance with system-related
and business-related constraints, which may vary from one
implementation to another. Moreover, it should be appreciated that
such a development effort might be complex and time consuming, but
would nevertheless be a routine undertaking of design, fabrication,
and manufacture for those of ordinary skill having the benefit of
this disclosure.
[0030] By measuring the artifact signal, the present technique
allows motion artifact to be separated from the plethysmographic
signal without the limiting assumptions of prior known techniques.
The present technique provides methods for measuring the motion
signal associated with changes in tissue optical properties and
using the measurement to compensate plethysmographic measurements
made at other wavelengths.
[0031] FIG. 1 is a block diagram of an exemplary pulse oximeter
that may be configured to implement the embodiments of the present
technique. The embodiments of the present technique can be a data
processing algorithm that is executed by the microprocessor 122,
described below. Light from light source 110 passes into patient
tissue 112, and is scattered and detected by photodetector 114. A
sensor 100 containing the light source and photodetector may also
contain an encoder 116 which provides signals indicative of the
wavelength of light source 110 to allow the oximeter to select
appropriate calibration coefficients for calculating oxygen
saturation. Encoder 116 may, for instance, be a resistor.
[0032] Sensor 100 is connected to a pulse oximeter 120. The
oximeter includes a microprocessor 122 connected to an internal bus
124. Also connected to the bus are a RAM memory 126 and a display
128. A time processing unit (TPU) 130 provides timing control
signals to light drive circuitry 132 which controls when light
source 110 is illuminated, and if multiple light sources are used,
the multiplexed timing for the different light sources. TPU 130
also controls the gating-in of signals from photodetector 114
through an amplifier 133 and a switching circuit 134. These signals
are sampled at the proper time, depending upon which of multiple
light sources is illuminated, if multiple light sources are used.
The received signal is passed through an amplifier 136, a low pass
filter 138, and an analog-to-digital converter 140. The digital
data is then stored in a queued serial module (QSM) 142, for later
downloading to RAM 126 as QSM 142 fills up. In one embodiment,
there may be multiple parallel paths of separate amplifiers,
filters and A/D converters for multiple light wavelengths or
spectra received.
[0033] Based on the value of the received signals corresponding to
the light received by photodetector 114, microprocessor 122 will
calculate the oxygen saturation using various algorithms. These
algorithms require coefficients, which may be empirically
determined, corresponding to, for example, the wavelengths of light
used. These are stored in a ROM 146. In one embodiment of a
two-wavelength system, the particular set of coefficients chosen
for any pair of wavelength spectra is determined by the value
indicated by encoder 116 corresponding to a particular light source
in a particular sensor 100. In one embodiment, multiple resistor
values may be assigned to select different sets of coefficients. In
another embodiment, the same resistors are used to select from
among the coefficients appropriate for an infrared source paired
with either a near red source or far red source. The selection
between whether the near red or far red set will be chosen can be
selected with a control input from control inputs 154. Control
inputs 154 may be, for instance, a switch on the pulse oximeter, a
keyboard, or a port providing instructions from a remote host
computer. Furthermore, any number of methods or algorithms may be
used to determine a patient's pulse rate, oxygen saturation or any
other desired physiological parameter. For example, the estimation
of oxygen saturation using modulation ratios is described in U.S.
Pat. No. 5,853,364, entitled "METHOD AND APPARATUS FOR ESTIMATING
PHYSIOLOGICAL PARAMETERS USING MODEL-BASED ADAPTIVE FILTERING,"
issued Dec. 29, 1998, and U.S. Pat. No. 4,911,167, entitled "METHOD
AND APPARATUS FOR DETECTING OPTICAL PULSES," issued Mar. 27, 1990,
both of which are incorporated herein by reference in their
entirety. Furthermore, the relationship between oxygen saturation
and modulation ratio is further described in U.S. Pat. No.
5,645,059, entitled "MEDICAL SENSOR WITH MODULATED ENCODING
SCHEME," issued Jul. 8, 1997 and incorporated herein by reference
in its entirety.
[0034] Having described an exemplary pulse oximeter above, the
methods for reducing noise, including motion artifact effects in
the received signals, according to embodiments of the present
technique, are described below.
[0035] FIG. 2 is a plot of the average plethysmographic amplitude
as a function of wavelength measured through the earlobe of 36
subjects, and normalized to measurements at a wavelength of
approximately 900 nm. Measurements, such as those shown in FIG. 2,
reveal that the amplitude of the photoplethysmographic waveform
diminishes as a function of wavelength between approximately 900
and 1300 nm, having a minimum value at approximately 1285 nm. The
inventors herein have discovered that at wavelengths beyond
approximately 900-920 nm, water, which is at much higher
concentrations than hemoglobin, also becomes a major light absorber
in tissue. FIG. 3 is a graph of some of the light absorbing
components found in blood at typical concentrations, in units of
absorbance in cm.sup.-1 vs. wavelength in nm. FIG. 3 shows that at
approximately 1300 nm, blood should have only about 20% as much
total absorbance as at 900 nm, with water being the dominant
absorber. This theoretical model is in rough agreement with the
pooled data shown in FIG. 2, where average plethysmographic
amplitude was about 1/3 as much at 1300 nm as at 900 nm.
[0036] FIG. 4 is a graph of absorption spectra (cm.sup.-1) of the
principal components in human skin, scaled to typical physiological
concentration, as a function of wavelength in nm. This figures
shows that the absorbance due to water has a peak value at
approximately 1180 nm, and that similar peaks are present for
protein at slightly above 1150 and for lipids at approximately 1200
mm.
[0037] FIG. 5 is a graph of absorption spectra of the principal
components in human skin, scaled to equal volume-fraction
concentration. This figure shows that at approximately 1185 nm, the
volume-fraction scaled absorbance for water, lipids and proteins
are approximately equal.
[0038] While not being limited to any particular theory, the
present inventors have, particularly in plethysmographic data from
reflectance sensors, noted a weaker effect of water than would be
theoretically predicted from absorption spectra. One potential
reason for this effect lies in the fact that hemoglobin is largely
confined to the blood vessels, whereas water is present at high
concentrations both in the blood vessels and in the surrounding
tissue. As a result, the pulse-induced expansion of arterial
vessels through a tissue bed results in a localized increase in
hemoglobin concentration, but only a small net change in water
concentration. To the extent that the water concentration in the
blood is equal to the water concentration in tissue, the change in
light absorption by water is expected to approach zero.
[0039] The embodiments of the present technique exploit the finding
that in spectral regions where hemoglobin absorbs weakly and water
absorbs strongly, the plethysmograph is more sensitive to
motion-related events that perturb tissue than arterial pulsation,
compared with spectral regions where hemoglobin is a strong
absorber and water is a weak absorber.
[0040] The weak magnitude of the plethysmograph in regions of
strong water absorption is exploited to enable the separation of
arterial-pulse-related signal from a motion artifact signal. By
measuring the optical plethysmograph at a wavelength where water is
the dominant absorber, the change in tissue optical properties
associated with motion or pressure can be measured, with little
interference from the underlying arterial pulsation.
Plethysmographs at four near-infrared wavelengths measured through
a human ear undergoing occasional motion are shown in FIG. 6, in
absorbance units vs. scaled time (i.e., time per point is 43 ms).
At approximately 920 nm, where hemoglobin absorption is strong and
water absorption is weak, the plethysmograph contains regular
arterial pulsations that are interrupted occasionally by
motion-related events. As the wavelength is increased to
approximately 1300 nm, where water is the predominant absorber, the
arterial pulsations diminish and the measured signal becomes
largely due to the motion-related events.
[0041] By combining the plethysmograph measured in a spectral
region where water is the dominant absorber with a plethysmograph
measured where blood is a major absorber, the motion-related signal
can be selectively removed. FIG. 7 shows the plethysmograph of a
human ear measured at approximately 920 nm, and the result of
subtracting a portion of the plethysmograph measured at
approximately 1180 nm from that measured at 920 nm. In particular,
FIG. 7 shows the plethysmograph of a human ear measured at 920 nm,
and the result of subtracting approximately 60% of the
plethysmograph measured at approximately 1180 nm from that measured
at approximately 920 nm. For different wavelength combinations,
other multipliers are used based on the ratios of the absorbance of
water as compared to that of oxy-hemoglobin or based on empirical
determination(s).
[0042] By applying the same analysis to a diverse pool of 36
patients measured in a hospital setting, an average signal to noise
increase of a factor of 2 of the plethysmograph at 910 nm was
observed. By allowing the multiplier for the 1180 nm plethysmograph
to vary between subjects, higher signal to noise improvements are
achieved.
[0043] Theoretical Model
[0044] The derivation below and the alternative description that
follows demonstrate mechanisms by which the effect of
motion-induced changes in optical scattering on a plethysmograph
measured at one wavelength can be compensated by plethysmographic
measurement at a second wavelength. These are provided as examples
of techniques for reducing motion-induced optical changes, but are
not the only mechanisms by which the present technique may
function, and thus are not meant to limit the embodiments of the
present technique.
[0045] A starting point for the analysis is the diffusion theory of
light transport in tissue (for example, see "Diffusion Theory of
Light Transport", Willem M. Star, in Optical-Thermal Response of
Laser-Irradiated Tissue, edited by Ashley J. Welch and Martin J. C.
van Gemert, Plenum Press, New York, 1995, pgs. 131-206). In the
case where the transport-corrected scattering coefficient,
.mu.'.sub.s, is much larger than the absorption coefficient,
.mu..sub.a, the diffuse intensity of light, I(.lamda.), measured at
wavelength, .lamda., by a detector positioned a distance, l, away
from a light source, can be described as follows (for example, see
"Measurement of Blood Hematocrit by Dual-Wavelength Near-IR
Photoplethysmography", Schmitt, J. M.; Guan-Xiong, G.; Miller, J.,
SPIE, Vol. 1641, 1992, pgs. 150-161): I(.lamda.).alpha. exp(-l
{square root over ((3 .mu..sub.a(.lamda.).mu..sub.s(.lamda.)))}
(eqn. 1)
[0046] For small changes in the absorption coefficient, such as
those caused by arterial pulsation, the resulting change in
intensity can be described by the derivative of intensity with
respect to the absorption coefficient: d I .function. ( .lamda. ) d
.mu. a .function. ( .lamda. ) I .function. ( .lamda. ) = AC
.function. ( .lamda. ) DC .function. ( .lamda. ) = - l .times. 3
.times. .mu. s ' .function. ( .lamda. ) 4 .times. .mu. a .function.
( .lamda. ) .times. .DELTA. .times. .times. V art .times. .mu. a
art .function. ( .lamda. ) ( eqn . .times. 2 ) ##EQU1##
[0047] where .DELTA.V.sup.art is the fractional volume change due
to arterial pulsation, .mu..sub.a.sup.art is the absorption
coefficient of the arterial blood under measurement, AC(.lamda.)
refers to the time varying portion of the optical signal and DC
(.lamda.) refers to the average or non-time varying portion of the
optical signal.
[0048] The arterial oxygen saturation, SpO.sub.2, is estimated if
the AC-DC ratio described by equation 2 is measured at two
wavelengths, .lamda..sub.1 and .lamda..sub.2, that are chosen so
that oxy- and deoxy-hemoglobin are readily differentiated (e.g.,
.lamda..sub.1.about.approximately 660 nm,
.lamda..sub.2..about..approximately 910 nm): R = AC .function. (
.lamda. 1 ) DC .function. ( .lamda. 1 ) AC .function. ( .lamda. 2 )
DC .function. ( .lamda. 2 ) = .OMEGA. 12 .times. .mu. a art
.function. ( .lamda. 1 ) .mu. a art .function. ( .lamda. 2 ) ( eqn
. .times. 3 .times. a ) ##EQU2## where: .OMEGA. 12 = .mu. s '
.function. ( .lamda. 1 ) .times. .mu. a .function. ( .lamda. 2 )
.mu. s ' .function. ( .lamda. 2 ) .times. .mu. a .function. (
.lamda. 1 ) ( eqn . .times. 3 .times. b ) ##EQU3## from which: Sp
.times. .times. O 2 = .mu. a HHb .function. ( .lamda. 1 ) - R
.times. .times. .OMEGA. 12 - 1 .times. .mu. a HHb .function. (
.lamda. 2 ) R .times. .times. .OMEGA. 12 - 1 .function. ( .mu. a O
2 .times. Hb .function. ( .lamda. 2 ) - .mu. a HHb .function. (
.lamda. 2 ) ) + .mu. a HHb .function. ( .lamda. 1 ) - .mu. a O 2
.times. Hb .function. ( .lamda. 1 ) ( eqn . .times. 3 .times. c )
##EQU4##
[0049] where .mu..sub.a.sup.HHb and .mu..sub.a.sup.O.sup.2.sup.Hb
are the respective absorption coefficients for deoxy- and
oxy-hemoglobin in arterial blood, and R is the ratio of the AC to
DC ratios.
[0050] The effect of small changes in the scattering coefficient,
such as may be brought about by compression of tissue or motion
artifact, are as set forth below by eqn. 4: d I .function. (
.lamda. ) d .mu. s ' I .function. ( .lamda. ) = AC .function. (
.lamda. ) DC .function. ( .lamda. ) = - l .times. 3 .times. .mu. a
.function. ( .lamda. ) 4 .times. .mu. s ' .function. ( .lamda. )
.times. .DELTA..mu. s ' .function. ( .lamda. ) ( eqn . .times. 4 )
##EQU5##
[0051] By measuring the AC-DC ratio at a third wavelength,
.lamda..sub.3, chosen so that the absorption due to hemoglobin is
weak but the absorption due to water is strong, the effect of the
motion-induced scattering change are removed from the AC-DC
measurement at .lamda..sub.2 by subtracting the scaled AC-DC
measurement at .lamda..sub.3. The resulting motion-corrected
plethysmograph, P, can be expressed as: P = AC .function. ( .lamda.
2 ) DC .function. ( .lamda. 2 ) - AC .function. ( .lamda. 3 ) DC
.function. ( .lamda. 3 ) .times. .OMEGA. 23 - 1 ( eqn . .times. 5
.times. a ) ##EQU6## where: .OMEGA. 23 = .mu. s ' .function. (
.lamda. 2 ) .times. .mu. .times. .times. a .function. ( .lamda. 3 )
.mu. s ' .function. ( .lamda. 3 ) .times. .mu. a .function. (
.lamda. 2 ) ( eqn . .times. 5 .times. b ) ##EQU7##
[0052] When the effects of arterial pulsation (equation 2) and
motion artifact (equation 4) are additive, equation 5 is expanded
as follows: P = - l .times. 3 .times. .mu. s ' .function. ( .lamda.
2 ) 4 .times. .mu. o .function. ( .lamda. 2 ) .times. .DELTA.
.times. .times. V art .times. .mu. a art .function. ( .lamda. 2 ) -
l .times. 3 .times. .mu. a .function. ( .lamda. 2 ) 4 .times. .mu.
s ' .function. ( .lamda. 2 ) .times. .DELTA..mu. s ' .function. (
.lamda. 2 ) + .OMEGA. 23 - 1 .function. [ l .times. 3 .times. .mu.
s 1 .function. ( .lamda. 3 ) 4 .times. .mu. a .function. ( .lamda.
3 ) .times. .DELTA..mu. a .function. ( .lamda. 3 ) + l .times. 3
.times. .mu. a .function. ( .lamda. 3 ) 4 .times. .mu. s '
.function. ( .lamda. 3 ) .times. .DELTA..mu. s ' .function. (
.lamda. 3 ) ] ( eqn . .times. 6 ) ##EQU8##
[0053] When water absorption dominates the absorption of light by
tissue at .lamda..sub.3, and the water concentration in the
arteries and surrounding tissue is nearly equal, .DELTA..mu..sub.a
(.lamda..sub.3) is approximately zero, and equation 6 simplifies
to: P = - l .times. 3 .times. .mu. s ' .function. ( .lamda. 2 ) 4
.times. .mu. a .function. ( .lamda. 2 ) .times. .DELTA. .times.
.times. V art .times. .mu. a art .function. ( .lamda. 2 ) ( eqn .
.times. 7 ) ##EQU9##
[0054] Equation 7 depends only on the effect of arterial pulsation
at .lamda..sub.2; the effect of the motion artifact has been
removed. In a similar manner the plethysmograph measured at
.lamda..sub.3 may be used to remove the motion effects from the
plethysmograph measured at .lamda..sub.1. The corrected
plethysmographs measured at .lamda..sub.1 and .lamda..sub.2 may
then be combined and used to estimate oxygen saturation, as
described, for example, by equation 3.
[0055] Several wavelengths in the range between approximately 900
and 1300 nm and more specifically in the range between
approximately 1150 and 1350 nm have been tested and found effective
at reducing motion-artifact from plethysmographs measured at
approximately 910 nm. Wavelengths at the longer wavelength side of
this range have the advantage of weaker absorbance of hemoglobin
compared to that of water (for example, see FIGS. 3 and 4).
However, wavelengths at the shorter end of this range have the
advantage of reduced variation with changing tissue composition. As
can be seen in FIG. 5, where the major components of tissue have
been normalized to equal volume fraction, water, lipid, and
non-hemoglobin protein all have approximately equal absorbance at
approximately 1185 nm. Therefore the absorbance of tissue at
approximately 1185 nm will vary little with changes in the relative
concentration of these principal components.
[0056] It is known that the detection of light beyond approximately
1100 nm cannot readily be accomplished with the silicon (Si)
detectors that are commonly employed in commercial oximeters. For
example, the detector used to collect the data displayed in FIGS.
2-7 employed Indium Gallium Arsenide (InGaAs) as the photosensitive
material. The most common type of InGaAs detectors are sensitive to
light between approximately 800 and 1700 nm. Therefore, in a pulse
oximeter designed in accordance with the embodiments of the present
technique, with the conventional wavelengths of 660 and 890 nm, in
addition to a new light source that emits at wavelengths that are
absorbed strongly by water (such as approximately 1180 nm or
approximately between 900-1400 nm), an additional detector(s) is
used. One such scheme employs two detectors, one Si and one InGaAs,
placed side-by-side. An alternative arrangement uses a collinear
("sandwich") detector containing separate Si and InGaAs layers,
such as those commercially available, for example, from the
Hamamatsu corporation. Yet another alternate arrangement uses two
Si detectors placed symmetrically on either side of an InGaAs
detector. Alternately, a germanium detector (Ge) is used as a
substitute for the InGaAs detector.
[0057] An Implementation of Motion Noise Reduction Technique
[0058] A practical technique by which the effect of motion-induced
changes in optical scattering on a plethysmograph measured at one
wavelength may be reduced by plethysmographic measurement at a
second wavelength is described below. This technique and the
derivation above should be considered examples, and are not the
only mechanisms by which the present technique may function. They
are not meant to limit the embodiments of the present
technique.
[0059] In one example, three wavelengths of light are used: a red
wavelength at 660 nm, a near infrared (NIR) wavelength at 890 nm,
and a NIR wavelength at 1300 nm. The first two wavelengths are both
chosen to be primarily absorbed by hemoglobin, and the third
wavelength is chosen to be primarily absorbed by water. After these
wavelengths of light from the light source 110 (See FIG. 1) are
passed through the tissue, the light is collected by a
photodetector 114 (See FIG. 1) generating plethysmographs at each
frequency.
[0060] Turning now to FIG. 8, in an exemplary embodiment, the red
plethysmograph 200, the NIR plethysmograph 202, and the NIR (water)
plethysmograph 204 are pre-processed (Block 210) prior to use. In
this step, the waveforms are converted to a natural logarithm, and
may by filtered to reduce noise, such as with a bandpass filter.
The preprocessed plethysmographs 218 are then mathematically
combined (Block 220) to identify periods of high motion noise and
to generate plethysmographs with reduced motion noise.
[0061] In one such embodiment, the preprocessed NIR plethysmograph
214 and the preprocessed NIR (water) plethysmograph 216 are used to
identify periods of high and/or low motion noise. This is performed
by calculating a ratio-of-ratios, R.sub.1300,890, between the
absorbances at the NIR wavelength (890 nm) and the water wavelength
(1300 nm) (See Eqn. 3a above for an example). The value of this
ratio is less than 1.0 for periods when there are little or no
motion artifacts, ranging from around 0.2 to 0.7 for most subjects.
In one embodiment, a default value of 0.4 may be selected for
initial use by an algorithm as described herein. In an exemplary
embodiment, R.sub.1300,890 is calculated on two periods: once using
three seconds of data for rapid detection of motion artifacts, and
once using fifteen seconds of data for use in adjusting the
combined weights, as discussed further below.
[0062] In one embodiment, a three step process is used to generate
plethysmographs with reduced motion noise. The first step is to
subtract fractions (F) of the preprocessed NIR (water)
plethysmograph 216 (Preprocessed.sub.1300) from the preprocessed
red plethysmograph 212 (Preprocessed.sub.660), and the preprocessed
NIR plethysmograph 214 (Preprocessed.sub.890), to generate
corrected waveforms:
Corrected.sub.890=Preprocessed.sub.890-F.sub.1300,890*Preprocessed.sub.13-
00 (eqn. 8)
Corrected.sub.660=Preprocessed.sub.660-F.sub.1300,660*Preprocessed.sub.13-
00 (eqn. 9)
[0063] In such an embodiment, the second step is to rescale the
corrected waveforms to preserve the ratio-of-ratios 222
(R.sub.660,890) between the absorbance signals at the red (660 nm)
and NIR (890 nm) wavelengths, so that the coefficients in eqn. 3b
will not need to change. This is performed by estimating the
fractions (C.sub.890 and C.sub.660) of the arterial pulse that were
cancelled in Corrected.sub.890 and Corrected.sub.660:
Rescaled.sub.890=Corrected.sub.890/(1.0-C.sub.890) (eqn. 10)
Rescaled.sub.660=Corrected.sub.660/(1.0-C.sub.660) (eqn. 11) where:
C.sub.890=R.sub.1300,890*F.sub.1300,890 (eqn. 12)
C.sub.660=R.sub.1300,890*F.sub.1300,660 (eqn. 13) R.sub.660,890 222
may be supplied by the two wavelength oximetry algorithm 230, which
calculates this value for determination of the oxygen saturation.
An alternative method for rescaling Corrected.sub.660 is to add a
percentage of Corrected.sub.890 to maintain a constant value for
R.sub.660,890:
Rescaled.sub.660=Corrected.sub.660+(F.sub.1300,660/F.sub.1300,890)*C.sub.-
890*Rescaled.sub.890 (eqn. 14)
[0064] In this embodiment, the third step in generating
plethysmographs with reduced motion noise is to adjust the
fractions, F.sub.1300,660 and F.sub.1300,890, of the NIR (water)
plethysmograph 216 (Preprocessed.sub.1300) subtracted from the
other two waveforms. Mathematical techniques may be selected that
minimize the power, standard deviation, or amplitude of the
resulting waveforms. Alternatively, techniques may be chosen that
minimize the skewness of the derivative of the rescaled waveforms,
or enhance some other recognized metric, or combination of metrics,
of signal quality. The techniques for adjusting the fractions,
F.sub.1300,660 and F.sub.1300,890 may be selected based on their
efficacy in reducing saturation or pulse rate errors in
representative sets of oximetry data that include motion
artifact.
[0065] An example of one technique for calculating F.sub.1300,660
and F.sub.1300,890, is to use the summations given below: F 1300 ,
890 = Preprocessed 890 , t .times. Preprocessed 1300 , t
Preprocessed 1300 , t 2 - R 1300 , 890 .function. ( 1.0 - C 890 )
.times. Rescaled 890 , t 2 Preprocessed 1300 , t 2 ( eqn . .times.
15 ) F 1300 , 660 = Preprocessed 660 , t .times. Preprocessed 1300
, t Preprocessed 1300 , t 2 - R 1300 , 890 .function. ( 1.0 - C 890
) .times. R 660 , 890 .times. Rescaled 890 , t 2 Preprocessed 1300
, t 2 ( eqn . .times. 16 ) ##EQU10## These summations may be
adequately represented by the approximations shown below: F 1300 ,
890 = .times. Preprocessed 890 , t .times. Preprocessed 1300 , t
Preprocessed 1300 , t 2 - ( eqn . .times. 17 ) F 1300 , 660 =
.times. Preprocessed 660 , t .times. Preprocessed 1300 , t
Preprocessed 1300 , t 2 - ( eqn . .times. 18 ) ##EQU11## In one
implementation using these summations, a value of 0.03 has been
found to work well for .epsilon.. Alternatively, these summations
may be approximated with infinite impulse response (IIR) filters.
The values for F.sub.1300,660 and F.sub.1300,890 typically range
from 0.6-0.9, and, in one embodiment, may be limited to range
between 0.5-1.0 with a default value of 0.7. As will be understood
by those skilled in the art, these constants may vary due to
factors such as wavelength selection or sensor site or
geometry.
[0066] As shown in FIG. 8, the plethysmographs which have been
adjusted to reduce the noise motion artifacts, Rescaled.sub.890 226
and Rescaled.sub.660 224, are then used in a two wavelength
algorithm 230 to calculate a value for oxygen saturation 232 and
pulse rate 234. In one embodiment, the two wavelength algorithm 230
may be similar to that described in U.S. Pat. No. 5,853,364, but
without the preprocessing that has already been done in block 210
of FIG. 8.
[0067] The improvements afforded by this technique are illustrated
in the graph shown in FIG. 9. For this test, an oxygen sensor was
attached to a test subject's ear lobe, which is highly susceptible
to motion artifacts. As a control, another sensor was attached to a
digit on the test subject. This second sensor was connected to a
oximeter using a standard two wavelength algorithm. In the graph,
the oxygen saturation 300 calculated from the preprocessed red
plethysmograph 212 (See FIG. 8) and the preprocessed NIR
plethysmograph 214 (See FIG. 8), using a standard two wavelength
algorithm, showed a significant drop during periods of motion, such
as nodding or shaking of the head. A control value 304 was
calculated from the sensor located on the digit and remained
steady. In contrast to the oxygen saturation 300 calculated from
the uncombined preprocessed plethysmographs 218 (See FIG. 8), the
oxygen saturation 302 calculated from the combined plethysmographs
228 (See FIG. 8) closely tracked the control. This is further
illustrated by the % modulation curves at the bottom of the graph
in FIG. 9. Prior to correction, the % modulation signal 306 shows
the motion noise to be far larger then the % modulation signal 308
after the technique above is used.
[0068] In a larger test, a test group of 10 subjects using the
standard two wavelength algorithm showed a pooled
root-mean-square-difference (RMSD) in oxygen saturation of 4.55%,
with some periods of 25% errors, between the moving sensor and a
non-moving control. In contrast, the same data processed by the
three wavelength algorithm discussed above showed a RMSD of 2.61%
for the pooled subjects.
[0069] The values calculated in the algorithm detailed above may be
used in a number of ways to display more accurate information to
the user, while minimizing the load on the processor. For example,
turning back to FIG. 8, during periods of very low motion
artifacts, such as where R.sub.1300,890<0.85, the calculation
above may be deactivated and the oxygen saturation 232 and pulse
rate 234 calculated using the uncombined data from the preprocessed
plethysmographs 218. Conversely, in such an embodiment, during
periods of high motion artifacts, the calculation may remain active
or may be activated and the oxygen saturation 232 and pulse rate
234 calculated using the combined data from the preprocessed
plethysmographs 218. Alternatively, the value for R.sub.1300,890
could be used to gradually interpolate between values calculated
from the preprocessed plethysmographs 218 and the combined
plethysmographs 228. This technique would be useful in cases where
the NIR (water) plethysmograph 204 was weaker, perhaps due to small
pulse amplitude or a thick sensor site. In this case, the NIR
(water) plethysmograph 204 would have a poor signal-to-noise ratio,
and using the combined plethysmographs 228 only during periods of
high motion artifacts would provide the most accurate
information.
[0070] Additional useful modifications could take advantage of the
extra signals provided by the technique. For example, additional
preprocessing filters may be implemented prior to the calculation
of the adjusted waveforms. In another example, various algorithms
in the oximeter, such as sensor off detection, may continue to use
the preprocessed plethysmographs 218, while the oxygen saturation
and pulse rate calculation use the combined plethysmographs
228.
[0071] In addition, an alternative wavelength selection to the
above-described augmentation to conventional pulse oximetry is an
all-NIR pulse oximeter. An example of an all NIR oximeter is an
oximeter employing light sources emitting at approximately 940,
1040, and 1180 nm used in conjunction with a single InGaAs
detector. In addition to the advantage of requiring only one
detector, the all-NIR implementation has advantages associated with
the optical properties of tissue. The accuracy of measurements made
using pulse oximetry depends, in part, on the extent to which the
paths traveled by the different colors of light are the same. The
mean path length and penetration depth of light at a particular
wavelength traveling through tissue is strongly affected by the
absorption and scattering coefficients of tissue at that
wavelength. In conventional pulse oximetry, in order to achieve the
same mean path length and penetration depth at two wavelengths, the
scattering and absorption coefficients at the two wavelengths need
to be matched. The scattering of light by tissue decreases rapidly
as a function of wavelength, with the result that the scattering
properties of tissue at approximately 940, 1040, and 1180 nm will
be more closely matched than the scattering properties of tissue at
a combination of both visible and NIR wavelengths such as
approximately 660, 890, and 1180 nm, for reasons discussed below.
The absorption properties of oxy- and deoxy-hemoglobin are such
that at high oxygen saturation values the net (i.e., combined
effects of oxy and deoxy) absorption coefficient due to hemoglobin
will be matched reasonably well at 660 nm and 940 nm. However, as
oxygen saturation values decrease, the high absorption coefficient
of deoxy-hemoglobin at approximately 660 nm will result in an
increasingly strong mismatch between the net absorption coefficient
of hemoglobin at approximately 660 and approximately 940 nm. The
net absorption coefficients of hemoglobin at approximately 940 and
approximately 1040 nm, will be more closely matched than at
approximately 660 and approximately 940 nm, over the full range of
measurable oxygen saturation values.
[0072] The choice of the wavelength used to measure the
motion-artifact signal depends partially on the need for matching
the optical path length to that of the signals to be corrected.
Beyond approximately 950 nm, the absorption coefficient of water,
protein, and non-hemoglobin protein, in addition to that of
hemoglobin needs to be considered in order to achieve close
matching of path lengths. Although about 1300 nm is a currently
preferred wavelength for measuring the motion-artifact signal,
other alternative wavelength values are also effective, for
example, wavelengths between approximately 1050 and 1400 nm and
between approximately 1500 and 1850 nm.
[0073] The embodiments of the present technique may be practiced by
placing the optical components directly at the tissue interface, or
alternatively, by transporting the light to and from the tissue
with fiber optics. The former implementation has the advantage of
more efficient delivery and collection of the light, whereas the
latter implementation has the advantages of being less costly. The
less costly solution is enabled by the fact that when employing
fiber optic delivery, the light sources and detectors can reside in
the monitor as opposed to the sensor, and considering that such
components may be more expensive that the fiber, this will result
in a less expensive device.
[0074] As will be understood by those skilled in the art, other
equivalent or alternative methods for the measurement of motion
artifact signal associated with changes in tissue optical
properties, and using the measurement to compensate
plethysmographic measurements made at other wavelengths, according
to the embodiments of the present technique can be envisioned
without departing from the essential characteristics thereof. For
example, a combination of visible and NIR or an all NIR wavelength
combination may be used to make the measurements. Moreover,
individuals skilled in the art of near-infrared spectroscopy would
recognize that additional terms can be added to the algorithms used
herein to incorporate reflectance measurements made at additional
wavelengths and thus improve accuracy further. Also, light sources
or light emission optics other then LED's including and not limited
to incandescent light and narrowband light sources appropriately
tuned to the desired wavelengths and associated light detection
optics may be placed near the tissue location or may be positioned
within a remote unit; and which deliver light to and receive light
from the tissue location via optical fibers. Additionally, sensor
arrangements functioning in a back-scattering or a reflection mode
to make optical measurements of reflectances, as well as other
embodiments, such as those working in a forward-scattering or a
transmission mode may be used to make these measurements.
[0075] While the invention may be susceptible to various
modifications and alternative forms, specific embodiments have been
shown by way of example in the drawings and have been described in
detail herein. However, it should be understood that the invention
is not intended to be limited to the particular forms disclosed.
Rather, the invention is to cover all modifications, equivalents,
and alternatives falling within the spirit and scope of the
invention as defined by the following appended claims.
* * * * *